Hemodynamic monitoring of the heart can provide valuable physiological information regarding the functional state of the myocardium, which is intimately related to its mechanical behavior. The quantitative measurement of blood flow, or cardiac output (CO), defined as the total blood volume pumped by the heart per unit time, is one of the most useful parameters in assessing cardiac capability. It reflects not only the functional state of the heart but also the response of the entire circulatory system to acute and chronic disease and the impact of therapeutic interventions. Basal CO is related to body size and varies from 4 to 7 liters per minute in adult humans.
However, CO is also one of the more difficult parameters to measure. The calculation of CO includes a determination of both the heart rate (HR) and the heart stroke volume (SV), from which CO can be estimated.
The HR can be determined in a number of ways, among which the phonocardiogram is considered to be the most accurate method. However, due to practical difficulties the phonocardiogram method is generally not employed for any continuous or long-term monitoring. In practice, HR is most typically determined by the electrocardiogram (ECG). The analog ECG signal typically displays electrocardial events as perturbations, usually referred to as waves. The heartbeat is most clearly reflected in the ECG signal as an R wave peak between a pair of adjoining Q and S wave valleys. The commonly used method of automatically identifying the QRS wave pulses is the threshold method, in which the rate of voltage change between consecutive data points of the ECG signal is monitored and compared with a threshold value, where slopes exceeding the threshold value are deemed to be associated with a portion of the QRS pulse. While this method commonly detects the interval between consecutive R waves successfully more than eighty percent of the time, it typically has difficulty in dealing with sources of irregular signal components such as pacemakers, muscle noise, and 60 Hz interference, as well as nearby T or P waves which may also be associated with significant slope changes. More significantly, CO monitoring cannot be accomplished with an ECG alone, as this measurement does not reflect the true pumping action of the heart.
Both invasive and non-invasive methods are available for measurement of the SV component of CO, or CO directly. Traditionally, CO has been measured invasively by using one of various indicator-dilution methods. The current accepted standard is thermodilution, where a chilled dextrose or saline solution is used as the indicator. A catheter incorporating a thermistor is inserted through the right atrium and right ventricle, and into the pulmonary artery. The solution is injected into the right atrium as a 5-10 mL bolus, where it mixes with venous blood, causing the blood to cool slightly. Changes in blood temperature occur over time as the solution bolus is washed out of the heart, then passes by the thermistor located at the tip of the catheter in the pulmonary artery, and result in a detectable temperature change in the blood flowing through the artery proportional to the relative volumes of solution bolus and blood. Blood temperature is measured to create a thermal dilution curve from which CO is derived.
Although it is the standard in clinical medicine, thermodilution has several disadvantages. Because of heat loss through the catheter wall, several 5 mL injections are required to obtain a consistent value for CO. If the CO is low, leading to a broad dilution curve, it is difficult to obtain an accurate value. Respiration-induced variations in the pulmonary arterial blood temperature also confound the dilution curve when it is of low amplitude. Risks from the procedure include line sepsis, bacterial endocarditis, large vein thrombosis, cardiac arrhythmia, pneumothroax, and increased mortality (Connors, et al., JAMA, 276:889-897, 1996). Thus, although the invasive methods are considered the most accurate, the risks associated with their use are often an unacceptable trade-off, for they require direct access to the arterial circulation. In addition, older invasive methods are not suitable for repetitive measurements and normally cannot be performed outside a hospital. Furthermore, invasive methods are very demanding in terms of time consumption and the need for skilled personnel.
As an option to invasive, intermittent methods, recent research has focused on obtaining CO from noninvasive, continuous methods. Emerging noninvasive techniques include the partial carbon dioxide rebreathing Fick technique, the echo-doppler technique and endotracheal impedance cardiography. A more developed technique is impedance cardiography, also known as thoracic electrical bioimpedance or impedance plethysmography. Impedance cardiography has been found to be a non-invasive method with the potential for monitoring the mechanical activity of the heart with virtually no risk. It can be conveniently handled by nursing and non-technical staff. It can usually tolerate moderate patient movement and can be left unattended for continuous and long-term monitoring. With this method, the SV is estimated; CO is then obtained as the product of SV times HR. For impedance cardiography, a constant current source (I), of approximately 50 to 100 kHz, is injected across the thorax. The resulting voltage (V) is used to estimate the impedance, assuming the impedance is purely resistive, as Z=V/I. Because the skin impedance is two to ten times the value of the underlying body tissue impedance, four, rather than two, electrodes are used in a configuration which eliminates the skin impedance from the impedance measurement.
U.S. Pat. No. 3,340,867, now RE 30,101, to Kubicek et al. discloses an impedance plethysmography system which employs four electrodes, two around the neck and two around the torso of a patient, to provide an impedance difference signal from the two center electrodes. The outermost pair of electrodes apply a small magnitude, high frequency alternating current to the patient while the inner pair of electrodes are used to sense voltage levels on the patient above and below the patient's heart. The impedances of the patient at each of the inner pair of electrodes could be determined from the sensed voltages and known applied currents.
According to Kubicek et al., SV is related to impedance (Z) as follows: EQU SV=R.(L/Z.sub.0).sup.2.(VET).(dZ/dt.sub.max)
where R is blood resistivity, L is the distance between the inner voltage sensing electrodes, Z.sub.0 is the mean thoracic impedance determined from the inner, voltage sensing electrodes, VET is the ventricular ejection time and dZ/dt.sub.max is the maximum negative slope change of the time-differentiated impedance signal, which is the time-differentiated difference between the impedances determined at the center two electrodes. The above equation is referred to as Kubicek's equation.
Kubicek's equation is based upon a parallel column model of the thorax in which it is assumed:
(1) the thorax is a cylinder, consisting of two electrically conducting tissue paths, also of cylindrical form with uniform cross-sectional areas and homogenous conducting materials, the first path being the blood with a relatively low resistivity and the second path being all other tissues with relatively high resistivities; PA1 (2) the relationship between the maximum impedance change and the stroke volume during the cardiac cycle is linear; PA1 (3) impedance measurements of the individual specific tissue volumes are not very useful in developing the model (the parallel columns model relies on the intact thoracic measurements); and PA1 (4) at 100 kHz frequency, a physiologically safe frequency, the relative thoracic impedance changes are at a maximum, the effects of polarization are negligible, and the reactive component of impedance is small, especially when compared to the real component, so that the reactance could be ignored in determining impedance with only a small error. PA1 (1) poor correlation of the methods and models with the results of the more accepted invasive techniques; and PA1 (2) a relatively high dependance on skilled technical operators.
Unfortunately, this technology was never widely accepted because of its poor correlation with invasive methods. A major source of inaccuracy was the requirement that the maximum negative slope change of the time-differentiated impedance be calculated. Derivatives are inherently noisy, and the early methods employed in calculating this slope change were unreliable.
More recently, Sramek improved the original Kubicek relationship between impedance and stroke volume (U.S. Pat. No. 4,450,527). Assuming that the thorax can be modeled as a frustocone, stroke volume is determined as: EQU SV=(L.sup.3 /4.25Z.sub.0).VET.(dZ/dt.sub.max)
This improved relationship was based on a better assumption of the shape of thorax, and Sramek's disclosure provided a method to minimize ventilation noise artifact. While probably only modestly effective, the averaging of the impedance enabled measurements to be acquired during normal breathing. (Until this innovation, measurements could only be made while the patient held his breath). Even with these improvements, the correlation with thermodilution remained weak.
In a recent study, the accuracy of Sramek's BoMed NCCOM3 impedance cardiograph was evaluated against Thermodilution in 19 cardiac patients ranging from 29 to 75 years. Initially after cardiac surgery, the correlation coefficient square, r.sup.2, was r.sup.2 =0.30 (p=0.002) between the two methods. This measure decreased to r.sup.2 =0.26 (p=0.004) 2 to 4 hours after surgery (Yakimets, et al., Heart Lung, 24:194-206, 1995). Therefore, approximately 30% of the variance from thermodilution can be accounted for using the Sramek equation. Thus, the Sramek model illustrates some improvement and accuracy over the Kubicek model but its major assumptions are still similar to those of the Kubicek model.
Despite its advantages, impedance cardiography has not been well accepted by clinicians for two primary reasons:
It is believed that poor correlation, the primary reason, can be traced back to a single source, namely the continuing inability to relate impedance cardiography and its mathematical model directly to cardiac physiology.
Ziang Wang's Ph.D. dissertation at Drexel University further improved system accuracy in impedance cardiography. Using time-frequency analysis, which is extremely resistant to noise artifact, the maximum negative slope change of the impedance derivative was calculated. Calculation of the HR was also improved using time-frequency analysis. This development was disclosed in two patents (U.S. Pat. Nos. 5,309,917 and 5,443,073) and is incorporated into the IQ System supplied by Renaissance Technologies.
In a study of 68 cardiac patients whose mean age was 45.+-.20 years, 842 simultaneous thermodilution and IQ System measurements were made in emergency rooms (n=36), operating/postanesthesia rooms (n=4), and surgical/medical intensive/coronary care units (n=38). The resulting correlation coefficient squared was r.sup.2 =0.74 (p&lt;0.01) between the two methods (Shoemaker, et al., Crit Care Med. 22:1907-1912, 1994). Therefore, approximately 75% of the variance from thermodilution can be accounted for using the Wang method.
According to Shoemaker, et al., the difference between thermodilution and impedance CO measurements diverged when the control baseline impedance was less than 20 ohms. Reduced baseline impedance occurs with increased fluid in the chest, associated with pulmonary edema and plural effusion. With increased intrathoracic fluids, the injected current bypassed normal thoracic structures, leading to inaccurate estimates of CO, compared to thermodilution measurements. Furthermore, in patients where high CO values were associated with tachycardia and cardiac dysrhythmias, impedance cardiography underestimated corresponding thermodilution values. Since the dZ/dt waveform reflects changes in impedance throughout the cardiac cycle, the pulsatile component of flow dominates in impedance CO calculations. However, the proportion of pulsatile systolic flow to the more constant diastolic flow is less in tachycardia and hyperdynamic states.
One system which has used a neural network to compensate for various shortcomings in measuring CO is described in U.S. Pat. No. 5,579,778. However, this system relies on invasive measurements of CO, and thus does not address the shortcomings of invasive monitoring generally.
It would be desirable to provide non-invasive monitoring to estimate stroke volumes, cardiac output and related cardiac function parameters which correlate more closely with the stroke volumes, cardiac outputs and the like determined by means of recognized, accepted invasive procedures, but which does not require of operators the technical skills required by current systems, thereby permitting relatively long-term monitoring of the patient's condition.
It would also be desirable to compensate for the apparent nonlinearity between invasive and non-invasive measurements of cardiac output.